This invention relates to magnetic resonance (MR) methods. More specifically, this invention relates to methods useful with MR imaging techniques to undo the effects on the residual transverse magnetization due to magnetic field gradient pulses used to encode spatial information thereinto.
By way of background, the nuclear magnetic resonance phenomenon occurs in atomic nuclei having an odd number of protons or neutrons. Due to the spin of the protons and the neutrons, each such nucleus exhibits a magnetic moment, such that, when a sample composed of such nuclei is placed in a static, homogeneous magnetic field, B.sub.o, a greater number of nuclear magnetic moments align with the field than against the field to produce a net macroscopic magnetization M (also referred to as longitudinal magnetization) in the direction of the field. If perturbed from this preferred alignment, under the influence of the magnetic field B.sub.o, magnetization M precesses about the axis of the field at a frequency which is dependent on the strength of the applied magnetic field and on the characteristics of the nuclei. The angular precession frequency, .omega., also referred to as the Larmor frequency, is given by the Larmor equation .omega.=.gamma.B, in which .gamma. is the gryomagnetic ratio which is constant for each MR isotope and wherein B is the magnetic field acting upon the nuclear spins. It will be thus apparent that the resonant frequency is dependent on the strength of the magnetic field in which the sample is positioned.
The orientation of magnetization M, in equilibrium directed along the magnetic field B.sub.o, may be perturbed by the application of a magnetic field oscillating at the Larmor frequency. Typically, such a magnetic field, designated B.sub.1, is applied in a direction orthogonal to the direction of the static magnetic field by means of a radio frequency (RF) pulse through RF coils connected to a radio-frequency-transmitting apparatus. The effect of field B.sub.1 is to rotate magnetization M about the direction of the B.sub.1 field. This may be best visualized if the motion of magnetization M due to the application of the RF pulses is considered in a Cartesian coordinate system which rotates at a frequency substantially equal to the resonant frequency about the main magnetic field B.sub.o in the same direction in which the magnetization M precesses (i.e., the rotating frame). In this case, the B.sub.o field is typically chosen to be directed in the positive direction of the Z-axis, which, in the rotating frame, is designated Z' to distinguish it from the fixed-coordinate system. Similarly, the X- and Y-axes of the rotating frame are designated X' and Y'. Bearing this in mind, the effect of an RF pulse, then, is to rotate magnetization M, for example, from its direction along the positive Z' axis toward the transverse plane defined by the X' and Y' axes. An RF pulse having sufficient magnitude and duration to rotate magnetization M into the transverse plane (i.e., 90.degree. from the direction of the B.sub.o field) is conveniently referred to as a 90.degree. RF pulse. Similarly, proper selection of magnitude and duration of an RF pulse will cause magnetization M to change direction from the positive Z' axis to the negative Z' axis. This kind of an RF pulse is referred to as a 180.degree. RF pulse, or for obvious reasons, as an inverting pulse. It should be noted that a 90.degree. or a 180.degree. RF pulse will rotate magnetization M through the corresponding number of degrees from any initial direction of magnetization M. It should be further noted that an MR signal will only be observed if magnetization M has a net transverse component (perpendicular to B.sub. o) in the transverse plane. Assuming an initial orientation of magnetization M in the direction of the B.sub.o field, a 90.degree. RF pulse produces maximum net transverse magnetization in the transverse plane since all of magnetization M is in that plane, while a perfect 180.degree. RF pulse does not produce any transverse magnetization. As will be discussed hereinafter, in practice, perfect 180.degree. RF pulses are difficult to achieve in all regions of the object lying within the field of an RF coil so that magnetization M is rotated by either more or less than 180.degree.. This can adversely affect image quality.
RF pulses may be selective or nonselective. Selective pulses are typically modulated to have a predetermined frequency content so as to excite nuclear spins situated in preselected regions of the sample having precession frequencies as predicted by the Larmor equation. The selective pulses are applied in the presence of localizing magnetic field gradients (discussed hereinbelow). Nonselective pulses generally affect all of the nuclear spins situated within the field of the RF pulse transmitter coil and are typically applied in the absence of localizing magnetic field gradients.
Upon cessation of the RF excitation, magnetization M due to the excited nuclear spins begins to return to equilibrium under the influence of the B.sub.o field. Any transverse component of the magnetization M will rotate about the Z-axis. As it does so, the magnetic flux intercepts the conductors of the RF coil and induces therein a voltage, termed the free induction decay (FID) MR signal. As is well known, 180.degree. RF pulses can be used to refocus the nuclear spins, following an FID, to produce spin-echo signals which are useful in MR investigational techniques. For many materials the return to equilibirum is governed by two exponential time constants associated with longitudinal and transverse magnetizations. The time constants characterize the rate of return to equilibrium of these magnetization components following the application of perturbing RF pulses. The first time constant is referred to as the spin-lattice relaxation time (T.sub.1) and is the constant for return of the longitudinal magnetization to return to its equilibrium value. Spin-spin relaxation time (T.sub.2) is the constant for the transverse magnetization to return to its equilibrium value of zero in a perfectly homogeneous field B.sub.o . In fields having inhomogeneities, the time constant for transverse magnetization is governed by a constant denoted T.sub.2 *, with T.sub.2 * being less than T.sub.2.
There remains to be considered the use of magnetic field gradients to encode spatial information (used to reconstruct images, for example) into MR signals. Typically, three such gradients are necessary: EQU G.sub.x (t)=.differential.B.sub.o /.differential.x, EQU G.sub.y (t)=.differential.B.sub.o /.differential.y, and EQU G.sub.z (t)=.differential.B.sub.o /.differential.z.
The G.sub.x, G.sub.y, and G.sub.z gradients are positionally constant throughout the imaging slice, but their magnitudes are typically time dependent. The magnetic fields associated with the gradients are denoted, respectively, b.sub.x, b.sub.y, and b.sub.z, wherein EQU b.sub.x =G.sub.x (t)x, EQU b.sub.y =G.sub.y (t)y, EQU b.sub.z =G.sub.z (t)z,
within the volume.
In the recent past, MR has been developed into an imaging modality utilized to obtain images of anatomical features of human patients, for example. Such images depicting nuclear spin distribution (typically protons associated with water in tissue), T.sub.1 and/or T.sub.2 relaxation parameters are believed to be of medical diagnostic value in determining the state of health of examined tissue. Imaging data for constructing MR images can be collected using one of many available techniques. Typically, such techniques comprise a pulse sequence made up of a plurality of sequentially implemented views. Each view includes at least an RF excitation pulse and a magnetic field gradient pulse to encode spatial information into the MR signal.
One known MR imaging technique which is particularly useful is of the Fourier transform (FT) type, a variant of which is frequently referred to as "spin warp." The spin-warp tecnique is discussed in an article entitled "Spin Warp NMR Imaging and Applications to Human Whole Body Imaging" by W.A. Adelstein et al, Physics in Medicine and Biology, Vol. 25, pp. 751-756 (1980). Briefly, the spin-warp technique, which will be discussed in greater detail hereinafter, employs a variable amplitude phase-encoding magnetic field gradient pulse prior to the acquisition of spin-echo signals to phase-encode spatial information in the direction of this gradient. In a two-dimensional implementation, spartial information is encoded in a direction orthogonal to the phase-encoding axis by observing the spin-echo signal in the presence of a magnetic field gradient in a direction orthogonal to that of the phase-encoding gradient. In a typical pulse sequence, the magnitude of the phase-encoding gradient pulse is incremented monotonically in the temporal sequence of views.
In some embodiments of FT MR pulse sequences, it is desirable to undo (reverse) the phase encoding accomplished in one view prior to applying phase-encoding gradient pulse in a temporally adjacent subsequent view. One case where this is useful is when the phase-encoding amplitude order is not monotonic. Other than monotonic order may be used, for example, to reduce periodic motion ghost artifacts, as disclosed and claimed in U.S. patent application Ser. No. 683,071 assigned to the same assignee as the present invention and which is incorporated herein by reference. It has been found that when a view employing a large amplitude phase-encoding gradient pulse follows a view in which a small amplitude phase-encoding gradient pulse is used, residual transverse magnetization resulting from the small phase-encoding pulse can corrupt the measurement in the large phase-encoding view. This can have a deleterious effect on image quality. The effects of residual magnetization will be discussed next in greater detail.
In any imaging pulse sequence, a dynamic equilibrium is set up governed by T.sub.1 and T.sub.2 relaxation relative to pulse sequence repetition time TR. Thus, if TR&lt;2T.sub.2, some transverse magnetization from the preceding view will still be present when the new view is to begin. This magnetization will retain the phasing given it in the preceding view. In particular, if the last phase-encoding gradient magnitude were small, the residual isochromats will be little dephased. During the new view sequence, a component of this residual magnetization will rephase along with the newly generated magnetization. The resulting spin echo will consist, then, of the two components. If the current phase-encoding gradient is of high magnitude, the "new" spin-echo component can be small enough that the left over component (from a large amplitude previous spin echo) may be comparable, or even dominant, when this happens, the data acquired for the new view is corrupted.
In the normal sequential advancement of phase encoding, the residual component is always much smaller that the new (primary) component because the difference in the primary component from one view of the next is slight, and relaxation effects thus deplete the left over magnetization component.
In non-sequential advancement, a large preceding component, even when depleted by relaxation, can be significant relative to the signal for a large phase encoding.
It is, therefore, a principal object of the invention to provide a method for reversing the effects on the residual transverse magnetization due to gradient pulses employed in each view of an MR pulse sequence, such that the magnetization is left in the same state after each view, i.e., with no memory of the particular gradient amplitude employed).